Using a magnetic resonance imaging (“MRI”) system, a substantially uniform main magnetic field is usually generated within an examination region. The main magnetic field polarizes the nuclear spin system of a subject (or a portion thereof) being imaged within the examination region. Magnetic resonance is excited in dipoles which align with the main magnetic field by transmitting radio frequency excitation signals into the examination region. Conventionally, radio frequency pulses that are transmitted via a radio frequency (“RF”) coil assembly move the dipoles out of alignment with the main magnetic field, and cause a macroscopic magnetic moment vector to precess around an axis parallel to the main magnetic field. Then, a corresponding radio frequency magnetic signal is generated. The RF magnetic resonance signal is received by the RF coil assembly, and from the received RF signals, an image representation can be reconstructed for display, and/or data obtained for other purposes.
Intrinsic signal-to-noise ratio (“SNR”) of the MRI system can be proportional to the static magnetic field. The MRI technology development is moving towards higher static magnetic fields (e.g., 7 T, 8 T, or 9.4 T). However, an increase in the static magnetic field can increase the resonance frequency. Such higher resonance frequency can result in a more significant coil loss that otherwise could be neglected in a low field (e.g., 1.5 T or below). This coil loss may includes a radiation loss and a conductor loss. Thus, the efficiency of conventional passive radio frequency (“RF”) coils is reduced, likely significantly, when the resonance frequency increases. Such reduction in efficiency of the conventional passive RF coils may compromise the higher SNR that is obtained by increasing the static magnetic field. One of the objects of the present invention is to address such efficiency reduction, e.g., using an active RF coil.
Previously, active devices have been used in active antenna and active filter designs. However, the use of the active devices not been significantly utilized in the conventional MRI coil designs. One of the reasons for such lack of utilization is because the sample loss (as opposed to coil loss) is dominant in common low field (1.5 T) RF detection. As the static magnetic field increases, the resonance frequency increases, and signal lost due to the coil becomes an issue, and therefore minimizing loss due to the coil is desirable.